The field of the invention is coherent imaging using vibratory energy, such as ultrasound and, in particular, systems and methods for shearwave dispersion ultrasound vibrometry (“SDUV”).
There are a number of modes in which ultrasound can be used to produce images of objects. For example, an ultrasound transmitter may be placed on one side of the object and sound transmitted through the object to an ultrasound receiver placed on the other side of the object. With transmission mode methods, an image may be produced in which the brightness of each pixel is a function of the amplitude of the ultrasound that reaches the receiver (“attenuation” mode), or the brightness of each pixel is a function of the time required for the sound to reach the receiver (“time-of-flight” or “speed of sound” mode). In the alternative, the receiver may be positioned on the same side of the object as the transmitter and an image may be produced in which the brightness of each pixel is a function of the amplitude or time-of-flight of the ultrasound reflected from the object back to the receiver (“reflection,” “backscatter,” or “echo” mode).
There are a number of well known backscatter methods for acquiring ultrasound data. In the so-called “A-mode” method, an ultrasound pulse is directed into the object by an ultrasound transducer and the amplitude of the reflected sound is recorded over a period of time. The amplitude of the echo signal is proportional to the scattering strength of the reflectors in the object and the time delay is proportional to the range of the reflectors from the transducer. In the so-called “B-mode” method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-mode method and their amplitude is used to modulate the brightness of pixels on a display. The location of the transducer and the time delay of the received echo signals locates the pixels to be illuminated. With the B-mode method, enough data are acquired from which a two-dimensional image of the reflectors can be reconstructed. Rather than physically moving the transducer over the subject to perform a scan it is more common to employ an array of transducer elements and electronically move an ultrasonic beam over a region in the subject.
The ultrasound transducer typically has a number of piezoelectric elements arranged in an array and driven with separate voltages (“apodizing”). By controlling the time delay, or phase, and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements (“transmission mode”) combine to produce a net ultrasonic wave focused at a selected point. By controlling the time delay and amplitude of the applied voltages, this focal point can be moved in a plane to scan the subject.
The same principles apply when the transducer is employed to receive the reflected sound (“receiver mode”). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delays, or phase shifts, and gains to the echo signal received by each transducer array element.
There are a number of electronic methods for performing a scan using a transducer having an array of separately operable elements. These methods include linear array systems and phased array systems.
A linear array system includes a transducer having a large number of elements disposed in a line. A small group of elements are energized to produce an ultrasonic beam that travels away from the transducer, perpendicular to its surface. The group of energized elements is translated along the length of the transducer during the scan to produce a corresponding series of beams that produce echo signals from a two-dimensional region in the subject. To focus each beam that is produced, the pulsing of the inner elements in each energized group is delayed with respect to the pulsing of the outer elements. The time delays determine the depth of focus which can be changed during scanning. The same delay factors are applied when receiving the echo signals to provide dynamic focusing during the receive mode.
A phased array system commonly employs so-called phased array sector scanning (“PASS”). Such a scan is comprised of a series of measurements in which all of the elements of a transducer array are used to transmit a steered ultrasonic beam. The system then switches to receive mode after a short time interval, and the reflected ultrasonic wave is received by all of the transducer elements. Typically, the transmission and reception are steered in the same direction, θ, during each measurement to acquire data from a series of points along a scan line. The receiver is dynamically focused at a succession of ranges, R, along the scan line as the reflected ultrasonic waves are received. A series of measurements are made at successive steering angles, θ, to scan a pie-shaped sector of the subject. The time required to conduct the entire scan is a function of the time required to make each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio. For example, a total of 128 scan lines may be acquired over a sector spanning 90 degrees, with each scan line being steered in increments of 0.70 degrees.
The same scanning methods may be used to acquire a three-dimensional image of the subject. The transducer in such case is a two-dimensional array of elements which steer a beam throughout a volume of interest or linearly scan a plurality of adjacent two-dimensional slices.
Recently, an ultrasound technique for measuring mechanical properties of tissues called shearwave dispersion ultrasound vibrometry (“SDUV”) was developed and described, for example, in co-pending U.S. patent application Ser. Nos. 10/956,461 and 11/536,330, which are herein incorporated by reference in their entirety. In SDUV, a focused ultrasound beam, operating within FDA safety limits, is applied to a subject to generate harmonic shear waves in a tissue of interest. The propagation speed of the induced shear wave is frequency dependent, or “dispersive,” and relates to the mechanical properties of the tissue of interest. Shear wave speeds at a number of frequencies are measured by pulse echo ultrasound and subsequently fit with a theoretical dispersion model to inversely solve for tissue elasticity and viscosity. These shear wave speeds are estimated from the phase of tissue vibration that is detected between two or more points with known distance along the shear wave propagation path.
One feature of the SDUV method is the use of a so-called “binary pushing pulse” that allows the operation of one single array ultrasound transducer for both motion excitation and the echo signal detection. The transducer focuses ultrasound at one location, the “vibration origin,” to vibrate the tissue of interest and then electronically steers its focus to another location, a “motion detection point,” for echo signal vibration detection. Instead of continuously vibrating the tissue of interest, the “pushing” ultrasound is turned on during a vibration time period to vibrate the tissue and turned off to provide a time window for the pulse echo motion detection. When the pushing pulse is off, a series of short ultrasound pulses is transmitted to the motion detection location and a corresponding series of echo signals is received and processed to determine the tissue vibration. This intermittent pulse sequencing strategy allows both the production of a shear wave and the monitoring of its propagation at the same time with a single array transducer.
Tissue mechanical properties such as elastic modulus, or stiffness, and viscosity are often related to the pathological state of the tissue. Palpation, an ancient diagnostic tool, is still widely used by physicians today to examine patients by touch. However, the reliability and specificity of palpation varies based on physicians' experience, and is a subjective tool. Moreover, if abnormal tissue is located deep, with respect to the skin surface, its detection by palpation is often difficult or impossible. It has been recognized that tissue shear moduli have high dynamic ranges in biological tissues, and that these moduli significantly change during a pathological process.
Recently, noninvasive methods have been developed to quantitatively measure both tissue shear elasticity and viscosity, simultaneously, using so-called “ultrasound vibrometry” techniques. One such method uses ultrasound harmonic vibration and pulse-echo ultrasound detection, and is described by Y. Zheng, et al., in “Detection of Tissue Harmonic Motion Induced by Ultrasonic Radiation Force Using Pulse-Echo Ultrasound and Kalman Filter,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 2007; 54:290-300. This method uses the ultrasound radiation force to induce a shear wave in a tissue region with a single frequency at a time and uses the pulse-echo ultrasound to detect the shear wave propagation. This method requires repetitive measurements for several different harmonics. It also requires simultaneous vibration and detection that is problematic for practical implementations.
Another such method is SDUV, which is referred to above and additionally described, for example, by S. Chen, et al., in “Shearwave Dispersion Ultrasound Vibrometry (SDUV) for Measuring Tissue Elasticity and Viscosity,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, 2009; 56:55-62. As noted, this method delivers focused ultrasound that generates an acoustic radiation force in a tissue region. The radiation force induces vibrations in the tissue, as well as a propagating shear wave. The vibration motion created in response to the shear wave propagation is detected using ultrasound-based pulse-echo measurement methods and appropriate signal processing techniques for motion detection. The motion amplitude and phase at a specified frequency are extracted using a signal processing technique such as Fourier analysis or fast Fourier transform, or a filter such as a Kalman filter, and used for shear wave speed estimation. The speed of the induced shear wave is measured by evaluating the phase shift of the shear wave at a given frequency over a propagation distance.
Measurements of shear wave speeds at multiple frequencies are then used to fit to a model to solve for the shear elasticity and viscosity, such as the Voigt model. The shear wave speed measurements can be fit with any viscoelastic model and are not restricted to the Voigt model.
The SDUV method has great potential to measure the viscoelastic material properties of stiffening liver tissue in fibrosis or cirrhosis, arterial stiffening due to atherosclerosis, myocardial stiffening due to dysfunction, and other applications.
The radiation force in SDUV is generated by an ultrasound transducer that transmits ultrasound waves in response to, for example, one of two kinds of signals. The first kind of signal produces an amplitude modulated ultrasound wave with a modulation frequency, ωm, and an ultrasound, or carrier, frequency, ωc. This type of ultrasound wave produces a force that has a dynamic component at ωm and 2·ωm, when large carrier amplitude modulation is utilized, and at 2·ωm when double sideband and suppressed carrier amplitude modulation is utilized. Using such method, a continuous vibration of the tissue is produced, and shear waves propagate outwards from the axis of the force, which can be detected at the same time. A typical range for the shear wave frequencies is 50-1000 Hz, while the ultrasound frequencies range from 1-10 MHz. In this case, two transducers are required, one for generating the radiation force with a single frequency component and another for the detection. The two transducers can be replaced by one transducer array, of which transducer elements are divided into two groups: one for vibration and another for detection. This process is repeated for several different vibration frequencies to evaluate the dispersion of the shear wave speed.
This method is not without its drawbacks. One exemplary drawback is that two transducers having two different center frequencies are required because the vibration and detection are operated at the same time. If the two transducers have similar center frequencies and bandwidth, there will be interference between the signals used for detection and vibration. Another exemplary drawback is that measurement time is prolonged because the responses of the vibration at several frequencies are measured separately.
The second kind of signal produces tone bursts of ultrasound energy. Using this second kind of signal, SDUV intermittently vibrates tissue and detects the vibration over a distance. In this case, a set of N tone bursts of length Tb are repeated at a period of Tp, which corresponds to a rate, fp, equal to 1/Tp. This method could be thought of as amplitude modulation with a square wave with a duty cycle of Tb·fp. While the first SDUV approach produces a radiation force that has a single frequency component, the second approach produces a radiation force at the frequency, fp, and its harmonics. For example, fp may be 100 Hz, and as a result, the shear wave harmonic components would be at 200 Hz, 300 Hz, 400 Hz, and so on. The harmonics in the shear wave spectrum should not be confused with so-called “harmonic imaging” common in medical ultrasound imaging, as, for example, the frequencies involved with harmonic imaging are in the megahertz range. The advantage of this tone burst method is that it allows the use of the same transducer array for both vibration and detection at different times. It also measures the tissue response of several harmonics at the same time.
Despite these benefits, this method is also not without its drawbacks. One exemplary drawback is that only a small percentage of received samples carry a significant amount of vibration signal because the induced vibration quickly dissipates in time, and because the period between two vibration pulses, or “push pulses,” is too long. Moreover, the duty cycle of the push pulse is very low and only a few motion detection samples will have significant displacement present. Another exemplary drawback with this method is that the desired harmonics are determined by the period of the pulse sequence; thus, the period must be long enough to produce harmonics in the hundreds of Hz. On the other hand, a significant number of samples must be acquired within one period to satisfy the Shannon sampling theorem and to meet the Nyquist criterion. For example, if the push pulse repetition frequency, fp, is 156.25 Hz and eight harmonics are desired, the pulse repetition frequency (“PRF”) of the pulse-echo measurements for the motion tracking has to be 2.5 kHz or higher. In this example, 16 detection pulses are transmitted between two consecutive push pulses; however, because the tissue vibration quickly damps over time, once the excitation pulse is complete only the first few detection pulses will carry meaningful information about the vibration.
There is a significant amount of interference in isolating the motion at a specific frequency. The vibration induced by the periodic pulses includes the fundamental repetition frequency, fp, and its entire harmonics. The decrease of the force amplitude of the higher harmonics can be relatively slow as the frequency increases. In the motion estimation process, the other frequency components interfere with the estimation of the selected frequency component. This interference also includes artifacts that can occur from aliasing of the sampled signal where frequency components above one-half the PRF can overlap into the frequency range of interest.
Weak vibrations at higher harmonics may cause unreliable estimations for tissue shear viscoelasticity. The amplitude of the force at the nth harmonic frequency is proportional to:sinc(Tb·n·fp)  Eqn. (1);
where sinc(x)=sin(πx)/πx is the so-called cardinal sine function. Thus, the amplitude of the higher frequency components of the acoustic force is significantly smaller than for the lower frequency components. In addition, the tissue tends to substantially attenuate the high frequency vibrations. These factors can make it difficult to measure the shear wave speed components at higher frequencies.
It would therefore be desirable to provide a method for shearwave dispersion ultrasound vibrometry (“SDUV”) that produces vibratory motion in a manner such that it can be sufficiently measured before it dissipates in time, produces vibratory motion that does not produce significant interference during measurements, and produces vibratory motion that can be tailored such that the power of higher frequency components can be independently adjusted to offset attenuative losses.